The present invention relates generally to endoluminal stent-graft devices suitable for percutaneous delivery into a body through anatomical passageways to treat injured or diseased areas of the body. More particularly, the present invention relates to a method of bonding microporous polytetrafluoroethylene (“PTFE”) coverings over a stent scaffold in a manner which maintains unbonded regions to act as slip planes or pockets to accommodate planar movement of stent elements. In one embodiment of the present invention bonded and unbonded regions are formed by means of a mandrel which has a pattern of either raised projections or recesses in its surface which are either synchronous or asynchronous, respectively, with stent elements.
The use of implantable vascular grafts comprised of PTFE is well known in the art. These grafts are typically used to replace or repair damaged or occluded blood vessels within the body. However, if such grafts are radially expanded within a blood vessel, they will exhibit some subsequent retraction. Further, such grafts usually require additional means for anchoring the graft within the blood vessel, such as sutures, clamps, or similarly functioning elements. To minimize the retraction and eliminate the requirement for additional attachment means, those skilled in the art have used stents, such as those presented by Palmaz in U.S. Pat. No. 4,733,665 and Gianturco in U.S. Pat. No. 4,580,568 which patents are herein incorporated by reference, either alone or in combination with PTFE grafts.
For example, the stent described by Palmaz in U.S. Pat. No. 4,733,665 can be used to repair an occluded blood vessel. The stent is introduced into the blood vessel via a balloon catheter, which is then positioned at the occluded site of the blood vessel. The balloon is then expanded thereby expanding the overlying stent to a diameter comparable to the diameter of an unoccluded blood vessel. The balloon catheter is then deflated and removed with the stent remaining seated within the blood vessel because the stent shows little or no radial retraction. Use of radially expandable stents in combination with a PTFE graft is disclosed in U.S. Pat. No. 5,078,726 to Kreamer. This reference teaches placing a pair of expandable stents within the interior ends of a prosthetic graft having a length that is sufficient to span the damaged section of a blood vessel. The stents are then expanded to secure the graft to the blood vessel wall via a friction fit.
Although stents and stent/graft combinations have been used to provide endovascular prostheses that are capable of maintaining their fit against blood vessel walls, other desirable features are lacking. For instance, features such as increased strength and durability of the prosthesis, as well as an inert, smooth, biocompatible blood flow surface on the luminal surface of the prosthesis and an inert, smooth biocompatible surface on the abluminal surface of the prosthesis, are advantageous characteristics of an implantable vascular graft. Some of those skilled in the art have recently addressed these desirable characteristics by producing strengthened and reinforced prostheses composed entirely of biocompatible grafts and graft layers.
For example, U.S. Pat. No. 5,048,065, issued to Weldon, et al. discloses a reinforced graft assembly comprising a biologic or biosynthetic graft component having a porous surface and a biologic or biosynthetic reinforcing sleeve which is concentrically fitted over the graft component. The reinforcing sleeve includes an internal layer, an intermediate layer, and an external layer, all of which comprise biocompatible fibers. The sleeve component functions to provide compliant reinforcement to the graft component. Further, U.S. Pat. No. 5,163,951, issued to Pinchuk, et al. describes a composite vascular graft having an inner component, an intermediate component, and an outer component. The inner and outer components are preferably formed of expanded PTFE while the intermediate component is formed of strands of biocompatible synthetic material having a melting point lower than the material which comprises the inner and outer components.
Another reinforced vascular prosthesis having enhanced compatibility and compliance is disclosed in U.S. Pat. No. 5,354,329, issued to Whalen. This patent discloses a non-pyrogenic vascular prosthesis comprising a multilamellar tubular member having an interior stratum, a unitary medial stratum, and an exterior stratum. The medial stratum forms an exclusionary boundary between the interior and exterior strata. One embodiment of this prosthesis is formed entirely of silicone rubber that comprises different characteristics for the different strata contained within the graft.
The prior art also includes grafts having increased strength and durability, which have been reinforced with stent-like members. For example, U.S. Pat. No. 4,731,073, issued to Robinson discloses an arterial graft prosthesis comprising a multi-layer graft having a helical reinforcement embedded within the wall of the graft. U.S. Pat. No. 4,969,896, issued to Shors describes an inner elastomeric biocompatible tube having a plurality of rib members spaced about the exterior surface of the inner tube, and a perforate flexible biocompatible wrap circumferentially disposed about, and attached to, the rib members.
Another example of a graft having reinforcing stent-like members is disclosed in U.S. Pat. No. 5,123,917, issued to Lee which describes an expandable intraluminal vascular graft having an inner flexible cylindrical tube, an outer flexible cylindrical tube concentrically enclosing the inner tube, and a plurality of separate scaffold members positioned between the inner and outer tubes. Further, U.S. Pat. No. 5,282,860, issued to Matsuno et al. discloses a multi-layer stent comprising an outer resin tube having at least one flap to provide an anchoring means, an inner fluorine-based resin tube and a mechanical reinforcing layer positioned between the inner and outer tubes.
Still another stent containing graft is described in U.S. Pat. No. 5,389,106 issued to Tower which discloses an impermeable expandable intravascular stent including a dispensable frame and an impermeable deformable membrane interconnecting portions of the frame to form an impermeable exterior wall. The membrane comprises a synthetic non-latex, non-vinyl polymer while the frame is comprised of a fine platinum wire. The membrane is attached to the frame by placing the frame on a mandrel, dipping the frame and the mandrel into a polymer and organic solvent solution, withdrawing the frame and mandrel from the solution, drying the frame and mandrel, and removing the mandrel from the polymer-coated frame.
Microporous expanded polytetrafluoroethylene (“ePTFE”) tubes may made by any of a number of well-known methods. Expanded PTFE is frequently produced by admixing particulate dry polytetrafluoroethylene resin with a liquid lubricant to form a viscous slurry. The mixture is poured into a mold, typically a cylindrical mold, and compressed to form a cylindrical billet. The billet is then ram extruded through an extrusion die into either tubular or sheet structures, termed extrudates in the art. The extrudates consist of extruded PTFE-lubricant mixture called “wet PTFE.” Wet PTFE has a microstructure of coalesced, coherent PTFE resin particles in a highly crystalline state. Following extrusion, the wet PTFE is heated to a temperature below the flash point of the lubricant to volatilize a major fraction of the lubricant from the PTFE extrudate. The resulting PTFE extrudate without a major fraction of lubricant is known in the art as dried PTFE. The dried PTFE is then either uniaxially, biaxially or radially expanded using appropriate mechanical apparatus known in the art. Expansion is typically carried out at an elevated temperature, e.g., above room temperature but below 327° C., the crystalline melt point of PTFE. Uniaxial, biaxial or radial expansion of the dried PTFE causes the coalesced, coherent PTFE resin to form fibrils emanating from nodes (regions of coalesced PTFE), with the fibrils oriented parallel to the axis of expansion. Once expanded, the dried PTFE is referred to as expanded PTFE (“ePTFE”) or microporous PTFE. The ePTFE is then transferred to an oven where it is sintered by being heated to a temperature above 327° C., the crystalline melt point of PTFE. During the sintering process the ePTFE is restrained against uniaxial, biaxial or radial contraction. Sintering causes at least a portion of the crystalline PTFE to change from a crystalline state to an amorphous state. The conversion from a highly crystalline structure to one having an increased amorphous content locks the node and fibril microstructure, as well as its orientation relative to the axis of expansion, and provides a dimensionally stable tubular or sheet material upon cooling. Prior to the sintering step, the lubricant must be removed because the sintering temperature of PTFE is greater than the flash point of commercially available lubricants.
Sintered ePTFE articles exhibit significant resistance to further uniaxial, or radial expansion. This property has lead many in the art to devise techniques which entail endoluminal delivery and placement of an ePTFE graft having a desired fixed diameter, followed by endoluminal delivery and placement of an endoluminal prosthesis, such as a stent or other fixation device, to frictionally engage the endoluminal prosthesis within the lumen of the anatomical passageway. The Kreamer Patent, U.S. Pat. No. 5,078,726, discussed above, exemplifies such use of an ePTFE prosthetic graft. Similarly, published International Applications No. WO95/05132 and WO95/05555, filed by W. L. Gore Associates, Inc., disclose balloon expandable prosthetic stents which have been covered on inner and outer surfaces by wrapping ePTFE sheet material about the balloon expandable prosthetic stent in its enlarged diameter, sintering the wrapped ePTFE sheet material to secure it about the stent, and crimping the assembly to a reduced diameter for endoluminal delivery. Once positioned endoluminally, the stent-graft combination is dilated to re-expand the stent to its enlarged diameter returning the ePTFE wrapping to its original diameter.
Thus, it is well known in the prior art to provide an ePTFE covering which is fabricated at the final desired endovascular diameter and is endoluminally delivered in a folded or crimped condition to reduce its delivery profile, then unfolded in vivo using either the spring tension of a self-expanding, thermally induced expanding structural support member or a balloon catheter. However, the known ePTFE covered endoluminal stents are often covered on only one surface of the stent, i.e., either the lumenal or abluminal wall surface of the stent. Where the stent is fully covered on both the luminal and abluminal wall surfaces of the stent, the covering completely surrounds the stent elements and fills the stent interstices. When the encapsulated stent is comprised of shape memory alloy, characteristics of the stent make it necessary to encapsulate in the “large” state and then compress the encapsulated stent for delivery. In this case encapsulation either increases the device's resistance to compression, or increases the delivery profile of the device as compression causes the polymeric material to fold or buckle around the stent. Perhaps the most serious problem is that the folding during compression actually encompasses folding of the stent itself, which unduly stresses the stent material and may result in structural failure.
In contrast to the prior art, the present invention provides a method to encapsulate a stent in ePTFE whereby the structure contains pockets or regions where the ePTFE layers are not adhered to one another allowing the stent to contract or expand without being encumbered by ePTFE and without folding or stressing the stent itself.
As used herein, the following terms have the following meanings:
“Fibril” refers to a strand of PTFE material that originates from one or more nodes and terminates at one or more nodes.
“Node” refers to the solid region within an ePTFE material at which fibrils originate and converge.
“Internodal Distance” or “IND” refers to a distance between two adjacent nodes measured along the longitudinal axis of fibrils between the facing surfaces of the adjacent nodes. IND is usually expressed in micrometers (μm).
“Node Length” as used herein refers to a distance measured along a straight line between the furthermost end points of a single node which line is perpendicular to the fibrils emanating from the node.
“Nodal Elongation” as used herein refers to expansion of PTFE nodes in the ePTFE microstructure along the Node Length.
“Longitudinal Surface” of a node as used herein refers to a nodal surface from which fibrils emanate.
“Node Width” as used herein refers to a distance measured along a straight line, drawn parallel to the fibrils, between opposing longitudinal surfaces of a node.
“Plastic Deformation” as used herein refers to the deformation of the ePTFE microstructure under the influence of a expansive force which deforms and increases the Node Length and results in elastic recoil of the ePTFE material less than about 25%.
“Radially Expandable” as used herein to describe the present invention refers to a property of the ePTFE tubular member to undergo radially oriented Plastic Deformation mediated by Nodal Elongation.
“Structural Integrity” as used herein to describe the present invention in terms of the ePTFE refers to a condition of the ePTFE microstructure both pre- and post-radial deformation in which the fibrils are substantially free of fractures or breaks and the ePTFE material is free of gross failures; when used to describe the entire device “Structural Integrity” may also include delamination of the ePTFE layers.
Endoluminal stent devices are typically categorized into two primary types: balloon expandable and self-expanding. Of the self-expanding types of endoluminal stent devices, there are two principle sub-categories: elastically self-expanding and thermally self-expanding. The balloon expandable stents are typically made of a ductile material, such as stainless steel tube, which has been machined to form a pattern of openings separated by stent elements. Radial expansion is achieved by applying a radially outwardly directed force to the lumen of a balloon expandable stent and deforming the stent beyond its elastic limit from a smaller initial diameter to an enlarged final diameter. In this process the slots deform into “diamond shapes.” Balloon expandable stents are typically radially and longitudinally rigid and have limited recoil after expansion. These stents have superior hoop strength against compressive forces but should this strength be overcome, the devices will deform and not recover.
Self-expanding stents, on the other hand, are fabricated from either spring metal or shape memory alloy wire which has been woven, wound or formed into a stent having interstices separated with wire stent elements. When compared to balloon-expandable stents; these devices have less hoop strength but their inherent resiliency allows them to recover once a compressive force that results in deformation is removed.
Covered endoluminal stents are known in the art. Heretofore, however, the stent covering has been made of a polymeric material which has completely subtended the stent interstices, that is, the stent was completely embedded in the polymeric material. This has posed difficulty particularly with the self-expanding stents. To preserve their self-expanding property, all covered self-expanding stents have been covered with a polymeric covering while the stent is in its unstrained dimensional condition (i.e., its native enlarged diameter). Yet in order to deliver a covered stent it must be constricted to a smaller delivery diameter. Radial compression of a stent necessarily causes the individual stent elements to traverse the stent interstices and pass into proximity to a laterally adjacent individual stent element, thereby occupying the previously open interstitial space. Any polymeric material which subtends or resides within the previously open interstitial space will necessarily be displaced, either through shearing, fracturing or otherwise responding to the narrowing of the interstitial space as the stent is compressed from its enlarged unstrained diameter to its strained reduced diameter. Because the struts of the stent are completely encapsulated, resistance of the polymer may cause folding or stressing of the struts during compression.
It was recognized, therefore, that a need has developed to provide an encapsulating covering for a stent which is permanently retained on the stent, substantially isolates the stent material from the body tissue forming the anatomical passageway or from matter within the anatomical passageway, and which permits the stent to deform without substantial interference from the covering material.
It is, therefore, a primary objective of the present invention to provide a method for encapsulating an endoluminal stent such that the encapsulating covering forms non-adhered regions which act as slip planes or pockets to permit the individual stent elements to traverse a substantial surface area of interstitial space between adjacent stent elements without resistance or interference from the encapsulating covering, thereby avoiding damage or stress to the stent elements.
It is a further object of the present invention to use the pockets between the bonded regions to contain and deliver therapeutic substances.
It is another objective of the present invention to provide an apparatus for applying to and selectively adhering sections of the encapsulating covering about the stent, and to provide a selectively adhered encapsulated covered stent-graft device.